JP6190467B2 - Optical fiber type biomedical sensor system and blood vessel insertion type distributed pressure measuring device - Google Patents

Description

  The present invention relates to a medical measurement system, and more particularly, by collecting and analyzing biological information such as blood pressure using a distributed optical fiber sensor system that measures temperature, pressure, and strain distribution using an optical fiber cable. The present invention relates to a fiber optic biodiagnostic sensor system that provides biological diagnostic information, and a blood vessel insertion type distributed pressure measuring device that measures blood pressure and the like.

Conventionally, in the measurement of pressure or flow velocity in PCI (percutaneous transluminal coronary angioplasty), the coronary flow reserve ratio (FFR) is used as an important diagnostic index. Here, PCI is a treatment method for ischemic heart disease, and is accumulated in arterial blood vessels composed of cells containing atheroma, lipids, calcium and various fibrous bonds, and dead bodies of cells. This is a treatment that expands the coronary arteries of the stenotic heart by increasing the blood flow. FFR is an index for knowing the degree of blood flow inhibition due to this stenotic lesion, and is an amount expressed as a ratio with respect to the normal value of the blood flow at the distal portion of the lesion. Specifically, the pressure (Pa) of the aorta (proximal part) of the stenotic lesion and the intracoronary pressure (Pd) of the distal part are measured and obtained from Pd / Pa.
This will be described in more detail with reference to FIG. FIG. 6A is a schematic diagram showing a stenotic lesion occurring in a blood vessel, and an arrow indicates a blood flow direction. FIG. 6B is a model diagram of the pressure change in the blood vessel corresponding to FIG. In FIG. 6B, the vertical axis indicates the maximum pulsation pressure in the blood vessel, and the horizontal axis indicates the optical fiber length (Lof). Here, the optical fiber length is the distance from the start point to the sensor tip position when the predetermined start point of the optical fiber is zero. If the position corresponding to the lesioned part Pc1 is the stenosis position S1, and the position corresponding to the lesioned part Pc2 is the stenosis position S2, the intravascular pressure (blood pressure) at these positions is the reference pressure P as shown in the figure. Gradually decrease from ₀ (corresponding to Pa above). Assuming that the pressure at the stenosis position S1 after the gradual decrease is P ₁ and the pressure at the stenosis position S2 is P ₂ , P ₁ / P ₀ and P ₂ / P ₀ are the coronary blood flow reserve amounts at the respective positions. It is called the ratio (FFR). The PCI is applied when this value is 0.75 or less. Usually, Pa is measured at the tip of a guiding catheter, and Pd is measured by a pressure sensor at the tip of a dedicated catheter called a pressure wire.
In this case, in order to determine the symptom that does not stop at a single location, such as multiple stenosis, or the physiological situation as information after the stent is attached, pressure or flow rate measurement in PCI is performed. There is a demand for obtaining a distribution instead of a single value.

  More specifically, in consideration of the size of the heart valve and the axial size of the coronary artery stenosis, a resolution of about 3 mm to 5 mm is desired for measuring the axial pressure distribution and the like. In addition, the probe diameter for measurement is better for the coronary artery (thin diameter) and the heart valve membrane to be thinner (for example, 0.4 mmφ or less), and has a suitable rigidity as a probe suitable for measurement, and A device having a mechanism for supporting an optical fiber is required. Considering the above contents, it is difficult to satisfy the fiber outer diameter specification of 0.4 mmφ or less when a plurality of fibers are used in order to fulfill this measurement purpose. Furthermore, in addition to measuring the pressure of the affected area, temperature, flow rate, etc. can be measured at the same time, and in order to measure without affecting the heartbeat, a multifunctional sensor such as the use of a sensor that does not use an electrical sensor is used. It is required to be.

Conventionally, there is an FBG (Fiber Bragg Grating) sensor as an optical fiber sensor used for such a purpose, but this sensor needs to process the FBG into the fiber. In addition, the function of this sensor was originally to measure temperature through elongation of the optical fiber or thermal deformation, and it was difficult to take out only pure pressure measurements, and in addition to the previous proposals It is necessary to install a pressure conversion mechanism at a location where the FBG senses pressure, and continuous pressure measurement is not possible.
In addition, in order to satisfy the specifications for multi-functionalization (multi-measuring function) that can measure items other than pressure, more than three fibers are required. There was an obstacle (see, for example, Patent Document 1).
In addition, since the fiber portion between the FBG and the FBG does not have a sensor function, it is basically difficult to measure a continuous signal (see, for example, Non-Patent Document 1).

  Furthermore, rapid multipoint and multiparameter measurements of temperature and pressure are possible by coating the FBG site with Zn metal deposition or the like, but the sensor sensitivity is not sufficient. In addition, a change in the shape of the catheter containing the measurement cable used for medical purposes causes distortion in the measurement cable. The frequency change due to this distortion is larger than the frequency change due to pressure, and the pressure signal and distortion are increased. Difficult to distinguish from signal.

  In order to improve the above point, there is a sensor system using a single probe. This is a sensor system for measuring pressure and flow rate using four micro electro mechanical systems (MEMS) optical sensors, but it requires an opening on the probe surface corresponding to the sensor mounting position for pressure measurement. Therefore, the measurement can be performed only at a limited number of points with a certain interval, which causes a great trouble in actual use. In addition, a plurality of optical sensors are required, which is disadvantageous in reducing the diameter.

  Further, as shown in FIG. 18, since the structure is complicated, such as requiring a plurality of openings (refer to cross section A1A1, cross section B1B1, and cross section C1C1 in FIG. 18) in the pressure sensing portion on the probe surface, There is a problem from the viewpoint of safety. Further, for the purpose of multi-functionality, when an electric sensor, which is another sensor, is used in combination, the influence on the cardiopulmonary system must be taken into consideration (for example, see Patent Document 2). ).

  In addition to the above, the conventional technique further has the following general problems in use in measuring the pressure in the blood vessel. The first point is that the number of measurement points is limited by the number of sensors because the measurement by the sensor is a single point measurement and not a distribution measurement. The second point requires measurement by a plurality of sensors as described above. However, since the number of measurement points is limited to several, the length of the blood vessel that can be measured without moving the probe is shortened. The third point is that it is difficult to measure vascular stenosis having a plurality of lesions because the probe diameter cannot be reduced to a certain extent because a plurality of sensors are required, and the sensor structure is complicated. about. The fourth point is that there is an influence due to variations in sensor sensitivity because a plurality of fibers having different sensitivities are basically used, and there is a variation in sensitivity even though the sensor is temporarily used or disposable. For this reason, all sensors need to be calibrated, which is not suitable for mass production and increases waste. From this point of view, it is considered practically difficult to use a sensor according to the prior art as a sensor for measuring pressure in a blood vessel.

International Publication WO2011 / 048509A1 International publication WO2013 / 061281A1 JP 2010-216877 A

Stephen Kreger, Alex Sang, Naman Grag and Julia Michel, “High-resolution distributed fiber-optic sensing for dynamic structural monitoring” SPIE Newsroom., 14 June 2013, DOI: 10.1117 / 2.1201305.004826 Robert O. Bonow et.al, Braunwald ’s Heart Disease: A Textbook of Cardiovascular Medicine ”ninth edition, ELSEVIER SAUNDERS, 2012 Japan Society of Mechanical Engineers, Mechanical Engineering Handbook Design Edition β8 Bioengineering, The Japan Society of Mechanical Engineers, 2007, pp. 65-75.

As described above, in order to perform a desired measurement in PCI (for example, to measure the coronary flow reserve ratio (FFR)), a plurality of fibers having multiple functions are required, and Although it is necessary to be able to measure the pressure separately, in the conventional technique as described above, the strain sensitivity of the inorganic glass-based optical fiber is much larger than the pressure sensitivity when measuring the pressure. The error is large and the pressure cannot be measured directly.
Further, if a plurality of sensors are used, measurement at multiple points is possible. However, since continuous measurement is difficult, data to be acquired leaks, and sensor sensitivity varies among the plurality of sensors. There is a possibility that the medical condition cannot be accurately grasped by these.
Furthermore, since the FBG is artificially manufactured in the optical fiber, there is a case where functions such as a sensing function inherent in the optical fiber cannot be used.

  In contrast to the above-described conventional techniques, a measurement method that uses an optical fiber as a sensor, simultaneously separates two or more physical quantities such as pressure and strain of a measured object, and measures its distribution as independent measurement parameters It has been known. This measurement method uses a Brillouin or Rayleigh scattering frequency change or phase change (see, for example, Patent Document 3), and utilizes the fact that an optical fiber responds to various physical quantities such as strain, temperature, and pressure as a sensor. Is. Furthermore, some optical fibers utilize the property of simultaneously reacting to strain, temperature, and pressure, which will be described in more detail below.

そして、さらに、Δε、ΔＴの初期値ε_０、Ｔ_０が与えられれば、上記で求めたΔε、ΔＴから、これらの値が、例えば光ファイバの長尺方向の位置に対して連続した値（である分布値）として測定されることから、所定の位置におけるε、Ｔの値が求まることになる。Specifically, there is one that measures the pressure, strain, and temperature at the same time by combining frequency changes of two scattered lights of Brillouin and Rayleigh scattered light, and is evaluated by the following formula, for example. That is, in the measurement method using the frequency shift Δν _{B of} Brillouin scattering and the frequency shift Δν _R of Rayleigh scattering in a hybrid, the relational expression of each frequency shift and strain, temperature, and pressure change is the sensitivity coefficient C _ij (i = 1 to 2 and j = 1 to 3), it is expressed as Expression (1) and Expression (2). In this equation, since the sensitivity of ΔP is usually small compared to the other two items (see Patent Document 3), here, it is abandoned to directly measure the pressure, and conversely, blood pressure is applied to the optical fiber. So that the expressions (1) and (2) are evaluated by the expressions simplified to the expressions (3) and (4), and the blood pressure is continuously converted into the strain of the optical fiber. (Refer to reference numeral 4 in FIG. 3), a method for measuring the pressure distribution with high accuracy is proposed. The present invention proposes a technique for specifying a pressure by measuring strain in an inorganic glass-based optical fiber. In the case of an organic optical fiber, the present invention directly relates to equations (1) and (2). It is also possible to measure the pressure.

Further, if initial values ε ₀ , T _{0 of} Δε, ΔT are given, these values are continuous values with respect to the position in the longitudinal direction of the optical fiber, for example, from Δε, ΔT determined above ( Therefore, the values of ε and T at a predetermined position are obtained.

More specifically, in multi-variable simultaneous measurement by such hybrid technology, the frequency shift Δν _R due to Rayleigh scattering and the frequency shift Δν _B due to Brillouin scattering are simultaneously measured, and in order to further improve the measurement accuracy, Rayleigh scattering is used, An error may be reduced by filtering the Brillouin scattering frequency shift Δν _B. A measurement example of the relationship between each frequency shift and pressure is shown in FIG. In this case, for example, when a TW-COTDR (Tunable Wavelength Coherent Optical Time Domain Reflectometry) method, a BOCDA (Brillouin Optical Correlation Domain Analysis) method, or the like is used as the hybrid measurement method, the distortion accuracy is 0.079 με and the temperature accuracy is 0.009. It is possible to achieve ° C. The measurement accuracy is guaranteed up to the value of the laser frequency error (for example, 12 MHz). Incidentally, this frequency measurement accuracy corresponds to the pressure accuracy 19 kPa (= 3 psi) when the change in strain and temperature is not ignored in the equations (1) and (2).

  In addition, when the optical fiber coating material (protective film, specifically, for example, PFA, that is, tetrafluoroethylene / perfluoroalkyl vinyl ether copolymer) is subjected to pressure, the axial effect of volume change is the axis of the optical fiber. When the coating material has a lower rigidity than glass because it affects the strain in the direction, the pressure sensitivity can be apparently increased by increasing the thickness by the coating (see FIG. 2). However, until the maximum blood vessel insertion requirement of about 0.4 mmφ, the apparent pressure sensitivity is almost equal to the data of 0.25 mmφ, and in fact, the apparent pressure sensitivity does not increase so much.

By the way, when the case of measuring the coronary flow reserve ratio (FFR) related to the present invention is examined, the pressure range to be measured is −4.0 to 40.0 kPa (−30 to 300 mmHg), heat per 1 ° C. The effect is ± 40.0 Pa / ° C. The accuracy is ± 0.1 kPa (= ± 1%) at −4.0 to 6.7 kPa (−30 to 50 mmHg), and ± 0 at −6.7 to 40.0 kPa (−50 to 300 mmHg). A value of about 3 kPa (= ± 3%) is required (see Non-Patent Document 2).
The outer diameter of the catheter that can be used is preferably 0.46 mmφ or less (see Patent Document 2).

  The accuracy of the pressure to be measured was compared with the pressure accuracy in the method of measuring pressure, strain and temperature simultaneously by the hybrid measurement method combining Brillouin and Rayleigh scattered light, which is planned to be applied in this application. In this case, since the effect of the optical fiber coating material cannot be expected so much, there is a possibility that the pressure measurement accuracy cannot be satisfied if the optical fiber as a sensor is in a normal shape.

  The present invention has been made to solve the above-described problems, and simultaneously measures measurement data in which two or more physical quantities such as strain obtained by changing the temperature and pressure of a measured object are mixed using an optical fiber. By analyzing this measurement data, two or more physical quantities of the object to be measured can be separated and measured as independent measurement parameters, and multifunctional measurement can be realized with a small number of optical fibers of two or less. It is intended to do.

The blood vessel insertion type distributed pressure measuring device according to the present invention is:
It is inserted into the blood vessel of a living body and measures the temperature and pressure distribution at a predetermined location of the measured object,
An outer layer that is deformed by external pressure and does not allow the object to be measured to enter the interior;
A single-mode optical fiber that is deformed by temperature and strain;
The single mode optical fiber is disposed so as to be in contact with the central axis portion or the outer peripheral portion, propagates the pressure applied to the outer layer, continuously converts the pressure distribution into the strain of the single mode optical fiber, and either fixed at a plurality of points in a part of the single-mode optical fiber, the elongation of the center axis of the part and proportionally said single mode optical fiber ambient pressure and contact at a plurality of locations of said single-mode optical fiber A structure to be converted, the outer diameter of which is 0.
A structure that supports the rigidity of the entire catheter of 4 mm or less, connects the catheter tip that is the guide portion of the catheter and the operation handle of the catheter, and is rotatable by the operation of the operation handle .
It is equipped with.

According to the blood vessel insertion type distributed pressure measuring apparatus of the present invention, even if there are a plurality of measurement sites, it is possible to measure accurately at a time. In addition, there is an effect that multifunctional measurement can be realized with a small number of optical fibers of two or less. In addition , measurement is possible even for blood vessels that are thinner than conventional ones, and there is no need to make holes around the measurement probe, so there is less risk of damaging the living body when acquiring data for biodiagnosis, and safer measurement Is possible. Furthermore, according to the optical fiber type biomedical sensor system equipped with the blood vessel insertion type distributed pressure measuring device of the present invention, a plurality of physical quantities of a measurement target are separated and measured continuously with high accuracy as independent measurement parameters. be able to.

It is a figure showing the example of a measurement of the frequency shift by Rayleigh scattering of an optical fiber, the frequency shift by Brillouin scattering, and a pressure. It is a figure showing the example of a measurement of the relationship between the frequency shift by Rayleigh scattering at the time of changing an optical fiber diameter, and a pressure. It is a figure which shows an example of the sensor system for optical fiber type biodiagnosis which concerns on Embodiment 1 of this invention. It is the flowchart which showed the measurement, analysis, and display procedure of the sensor system for optical fiber type biodiagnosis concerning Embodiment 1 of the present invention, and the outline. It is a flowchart in the case of performing FFR determination. It is a model figure which shows the stenosis lesion part of a blood vessel, and a model figure of the pressure change in this stenosis lesion position. It is explanatory drawing which shows the sensor part used as the basis of the sensor system for optical fiber type biodiagnosis based on Embodiment 1 of this invention. It is explanatory drawing of the sensitivity evaluation model of the optical fiber structure material which concerns on Embodiment 1 of this invention. It is explanatory drawing which shows an example of the optical fiber structure material of the sensor part of the optical fiber type biodiagnosis sensor system which concerns on Embodiment 1 of this invention. It is a figure which shows an example of the optical fiber type biomedical sensor system which concerns on Embodiment 2 of this invention. It is explanatory drawing of the sensor part of the sensor system for optical fiber type biodiagnosis concerning Embodiment 2 of this invention. It is the flowchart which showed the measurement, analysis, and display procedure of the optical fiber type biodiagnostic sensor system which concerns on Embodiment 2 of this invention. It is explanatory drawing of the other sensor part of the optical fiber type biopsy sensor system which concerns on Embodiment 2 of this invention. It is a figure which shows an example of the sensor system for optical fiber type biodiagnosis concerning Embodiment 3 of this invention. It is a model figure for demonstrating the pulsation of the heart. It is explanatory drawing of the other optical fiber structure material of the optical fiber type biomedical sensor system of Embodiment 1-3 of this invention. It is explanatory drawing of the other optical fiber structural material of the sensor system for optical fiber type biodiagnosis of Embodiment 1-3 of invention. It is a figure which shows an example of the sensor part used for the conventional sensor system for biodiagnosis.

  Hereinafter, the optical fiber type biomedical sensor system of the present invention will be described with reference to the drawings. In addition, what attached | subjected the same code | symbol in each figure shows that it is the same, The description is abbreviate | omitted.

Embodiment 1 FIG.
First, the optical fiber type biodiagnostic sensor system according to the first embodiment of the present invention will be described with reference to FIG. The system shown in the first embodiment is a basic system of the sensor system for optical fiber biodiagnosis of the present invention.
This sensor system is mounted inside a blood vessel insertion type distributed pressure measuring device (hereinafter referred to as “catheter” in the present invention) 1 which is a component of the system inserted into a blood vessel of a living body to be measured. The sensor output obtained from the optical fiber that requires the sensor function is stored as data serving as a basis for performing a biological diagnosis, and this data is analyzed and displayed.
In the figure, the tip of the catheter 1 is equipped with a jet tip (hereinafter referred to as “J tip”) 2 which is a guide portion for facilitating insertion of the catheter into a blood vessel. Is inclined with respect to the axis of the catheter, and is connected to a metal structure 4 such as SUS at the rear end thereof. In some cases, in the outer layer portion corresponding to the outer periphery of the sensor tip portion connected to the rear end portion of the J tip 2, a spiral wire may be used instead of the rubber outer layer (described in detail later). ). Further, the surface of the catheter 1 has an outer layer 5 made of rubber or a soft skin material that does not adversely affect the human body, and blood to be measured does not pass through the outer layer 5, but the blood pressure passes through the soft outer layer 5. It is transmitted to the structure 4. The structure 4 has a function of converting a loaded pressure into an extension in the longitudinal direction, and a single mode optical fiber (hereinafter abbreviated as an SM fiber) 3 having a sensor function inside thereof is added to the longitudinal direction described above. Cause deformation of elongation. The catheter 1 is provided with an operation handle 6 for operating a main movement of the catheter at a hand portion. Although the SM fiber 3 is not in direct contact with blood, since the elongation of the structure 4 is proportional to the blood pressure, the axial strain of the SM fiber 3 is proportional to or converted to the pressure of the blood to be measured. Can do. In some cases, a jacket (not shown) for guiding the insertion of the catheter 1 into the living body on the surface of the living body at the insertion portion into the living body is provided on the outer peripheral portion of the catheter 1.
Actually, in a measurement site or a sick patient, the temperature of the entire blood vessel is usually not constant, and a measurement method that takes temperature non-uniformity into consideration is required. That is, when measuring pressure, it must also be possible to measure temperature.
The sensor output detected by the SM fiber 3 passes through the multi-core SM fiber 7 connecting the catheter 1 and the measuring device 8, and the measuring device 8 uses the Brillouin scattering frequency shift Δν _B and the Rayleigh scattering frequency shift Δν _R. (The signal is synchronized with the synchronization signal as necessary) and output to the data memory 9 as distribution data such as the temperature of blood in the blood vessel and the converted blood pressure. Thereafter, the data stored in the data memory 9 is analyzed by the analysis / display device 10 to obtain a coronary blood flow reserve ratio (FFR) or the like as diagnostic data as desired data, and the blood pressure distribution at the site to be diagnosed The data is output as signal data, and is displayed as a chart on a display or the like provided in the analysis / display device 10 as necessary.

A summary of the above measurement procedure is shown in FIG. In this figure, the measurement start time (measurement start time from measuring 8) and t _0, at Δt time interval SM fiber 3, the measurement is repeated. In each measurement, the frequency shift Δν _{R of} Rayleigh scattering and the frequency shift Δν _{B of} Brillouin scattering, which are sensor outputs, are measured.
Next, using these two frequency shifts Δν _R and Δν _B and the temperature T ₀ and the displacement ε ₀ in the reference state previously input to the measuring machine 8 as initial values, the temperature displacement ΔT and the strain change A distribution of Δε with respect to the measurement position is obtained (see Expressions (3) and (4)).
Next, a desired blood pressure (pressure) P is obtained using the measuring device 8 as a conversion by a function of the strain ε. Then, the pressure P obtained by the measuring device 8 is output to the data memory 9 and stored. Thereafter, the analysis / display device 10 displays the data analyzed in the data memory 9 as a distribution of the pressure P at time t in a desired display format. In this case, the time interval Δt is required to have a speed at which blood pulsations can be distinguished, and blood pressure distribution can be obtained at each time interval with reference to 100 times or more per second (in FIG. 15, 250 times per second). Example). Then, a coronary flow reserve ratio (FFR) or the like is obtained using the distribution when the pulsation pressure is maximum.

この値を測定対象となるすべての位置で求める（ＳＴ２）。
次に、ステップ３（ＳＴ３と略記する）では、計測値Ｐmes(ｚ、ｔ)から、ＳＴ２で得た中央値Ｐmed（ｚ）またはＰav（ｚ）を差し引き、差し引いた後の圧力の値をＰ（ｚ，ｔ）とおく（それぞれ式（７）、式（８）参照）。この処理により、カテーテル挿入時の静ひずみは除かれる。

最後にステップ４（ＳＴ４と略記する）で、測定位置ｚを変えて各位置での圧力Ｐ（ｚ，ｔ）の値を求め、ｚを横軸（距離）として、圧力の変化をプロットし、この変化から、ＦＦＲを求めて判定を行う。 Therefore, the procedure for determining the coronary blood flow reserve ratio (FFR) will be described in more detail with reference to FIG.
As shown in FIG. 4, since the pressure P is obtained as a function of the position (Lof) and the time t, if this position coordinate is set to z, the measured pressure P can be expressed as P = Pmes (z, t). The measured pressure Pmes (z, t) may include the influence of static strain generated when the catheter 1 is inserted. Therefore, a method for removing the influence of the static strain using the pressure measurement values obtained at the two positions A and B will be described with reference to FIG.
In step 1 of FIG. 5 (abbreviated as ST1), the pressure Pmes (z, t) measured at the position _A is represented by the symbol Z _A , and the pressure Pmes (z, t) measured at the position B is represented by the symbol Z _B. (In ST1, the vertical axis indicates pressure, and the horizontal axis indicates the time elapsed after the start of measurement). The measurement waveform of ST1 is a model diagram showing changes in left ventricular pressure corresponding to two heart beats. Next, in step 2 (abbreviated as ST2), a median value at the position A or the position B is obtained from the measured pressure Pmes (z, t) according to the equation (5) (see the equation (5)), or An average value at the position A or the position B is obtained according to the equation (6) (see the equation (6). The integral symbol ∫ is used).

This value is obtained at all positions to be measured (ST2).
Next, in step 3 (abbreviated as ST3), the median value Pmed (z) or Pav (z) obtained in ST2 is subtracted from the measured value Pmes (z, t), and the pressure value after subtraction is set to P (Z, t) is set (see equations (7) and (8), respectively). By this process, static strain at the time of catheter insertion is eliminated.

Finally, in step 4 (abbreviated as ST4), the measurement position z is changed to obtain the value of the pressure P (z, t) at each position, and the change in pressure is plotted with z as the horizontal axis (distance). From this change, the FFR is obtained and determined.

As described above, in the first embodiment, the blood pressure in the blood vessel of the living body to be measured is determined in a predetermined manner, both temporally and spatially, by using frequency shift data by two types of scattered light. Since the measurement can be continuously performed at the position, the physical quantity to be measured can be measured by one measurement only by using one SM fiber 3. Further, in the conventional point measurement sensor, measurement is difficult in the case of a blood vessel having a plurality of vascular stenosis due to the structure of the measurement probe. However, according to the present invention, the blood vessel can be continuously measured. Even when there are a plurality of stenosis, it is possible to cope (see FIG. 6). In FIG. 6, the coronary flow reserve ratio (FFR) is expressed as P ₁ / P ₀ or P ₂ / P ₀ (where P ₀ , P ₁ , and P ₂ are the same as those in ST1 to ST4). Use the value obtained along the line). FIG. 6 shows a case where there are a plurality of vascular stenosis. However, the present invention is not limited to this. Of course, measurement is possible even when there is only one vascular stenosis. Even in this case, the stenosis position can be obtained from the pressure measurement, which is considered useful as data for biodiagnosis.

The sensor portion used in this case will be described in more detail below. FIG. 7 shows a cross-sectional view of the structure of a general sensor portion in this case. In this figure, a J tip 2 made of plastic (for example, vinyl chloride) having an inclined portion is provided at the tip, and an SM fiber 3 is provided at the central axis portion of the sensor portion connected thereto. A structure 4 for transmitting external pressure to the outer periphery of the SM fiber 3 and converting the strain into strain of the SM fiber 3 is provided on the outer periphery thereof, and an outer layer 5 made of, for example, rubber is provided on the outer periphery thereof (FIG. 7). (A)).
A spiral wire layer 16 may be provided on the outer periphery of the outer layer to facilitate insertion of the catheter into the living body (FIG. 7B). Even if the wire layer 16 is attached, the state in which pressure is applied to the outer layer 5 does not change because blood passes between the wires.
The outer diameter of the SM fiber 3 shown in this figure is about 80 to 250 μm, and the structure 4 that is a protective layer of the SM fiber disposed on the outer periphery converts the applied pressure into strain, It is composed of SUS, and has a function of proportionally converting the ambient pressure into the amount of elongation in the longitudinal direction that is the axial direction of the SM fiber 3 (which will be described in more detail below). With this function, the strain of the SM fiber can be made proportional to the pressure. Moreover, the outer diameter as a catheter can be about 0.4 mm or less. For this reason, the specification as a biodiagnosis sensor to which the present invention is scheduled to be applied can be satisfied, and it can be effectively used in the measurement of pressure or flow velocity with PCI. Further, even when the blood vessel of the living body to be measured is a blood vessel having a relatively small outer diameter, pressure (blood pressure) measurement can be performed.

It is also conceivable that the structure 4 is integrated as a SM fiber coating. There is a possibility that the sensitivity received by the SM fiber is increased by changing the pressure received by the coating material.
In the case of increasing the sensitivity, as described above, when the coating material is lower in rigidity than glass, the pressure sensitivity can be apparently increased by increasing the thickness by the coating (see FIG. 2). However, in the field of measuring pressure or flow rate in PCI, to which the present invention is scheduled to be applied, the outer diameter of the catheter needs to be about 0.4 mm or less as described above. We cannot expect pressure sensitivity to increase several times or more. It is also effective to construct the SM fiber itself with a plastic material having low rigidity.
Therefore, in the following, a description will be given of an appropriate structural material that has a continuous, high-precision, uniform and simple structure and converts distributed pressure into continuous strain.

Therefore, first, in order to perform sensitivity evaluation by using an appropriate structure material that converts such pressure into strain, a model shown in FIG. 8 is used. In FIG. 8A, a model is considered in which an opening having a diameter D is formed on the outer peripheral surface of a cylindrical measuring probe in which an SM fiber is disposed on the central axis, and an external pressure acts on the opening. Here, the measurement probe simulates a catheter, and the external pressure corresponds to the blood pressure at the opening.
Normally, the structure is an axial object, but in this case, as shown in FIG. 8 (b), two halves arranged at an angle (angle θ) with respect to the SM fiber, which is displayed only in a half portion, are displayed. It is assumed that an elastic bar having the same shape receives external pressure (pressure P) acting on the opening on the outer periphery of the measurement probe at this opening. In other words, it is assumed that the structure is disposed at a distance h from the outer periphery of the measurement probe to the SM fiber and a distance (pitch) L between AB in the direction along the SM fiber axis between the two structures. In this case, as shown in FIG. 8C, the pressure P acting in the vertical direction (corresponding to the blood pressure in this portion) P is separated into two identical rods arranged obliquely at an angle θ with the horizontal axis. It can be considered as a model in which an equivalent axial force Tc acts in the horizontal direction (the axial direction of the SM fiber). This axial force Tc causes elongation and strain of the SM fiber, and the strain becomes a measurement target.

Therefore, the above model is evaluated by applying specific numerical values. Since the area of the opening is πD ^2/4, when the blood pressure at the opening portion and Pb, external pressure (pressure P) is expressed by ^{P = Pb × (πD 2/} 4). Therefore, when Pb = 100 (N / m ² ) and D = 0.25 mm, P = 100 (N / m ² ) × 4.9 × 10 ⁻⁸ (m ² ) = 4.9 × 10 ⁻⁶ ( N). When the pitch L = 10 mm and h = 0.2 mm, tan θ = 0.2 / 5 = 0.04. Therefore, assuming that θ << 1, θ≈tan θ≈sin θ. Therefore, Tc = P / 2sin θ = 4.9 × 10 ⁻⁶ (N) / (2 × 0.04) ≈6.13 × 10 ⁻⁵ (N). In a standard optical fiber, the strain with respect to the axial force of 1N is 1020 με, so the strain ε in this case is ε = 6.25 × 10 ⁻² (με). Empirically, it is considered that a strain of about 1 με is necessary, and it is necessary to employ a probe structure (catheter structure) that satisfies this.

As the above-described appropriate optical fiber structure material, for example, a frame structure made of a SUS material shown in FIG. 9 is proposed. This figure shows a structure 4 that converts the pressure for realizing the above contents into strain of an optical fiber (SM fiber in this case), and the SUS frame 11 is attached to the SM fiber 3 at both end portions (fixed portions) of the pitch L. It is fixed with an adhesive (see C1C1 portion in FIG. 9A, cross section C1C1 in FIG. 9B). The SUS frame 11 has a structure bulging outward from the SM fiber axis by the size indicated by h. Note that the upper and lower envelopes of the SUS frame 11 in FIG. 9 (a) indicate the maximum diameter of the swollen portion, and the outer layer 5 that is the propagation portion of the external pressure of the sensor shown in FIG. 9 (c). It is also a boundary line.
Further, as shown in FIG. 9B (the figure on the left side of FIG. 9B), the SM fiber 3 is bonded to the outer periphery of the SM fiber 3 with an adhesive 12 at a fixed portion (the portion shown by the cross section C1C1). The formed SUS frame 11 is concentrically bonded and fixed. The SM fiber 3 and the SUS frame 11 are configured so as to be concentrically separated from each other as shown as a typical example in a cross section D1D1 (right side view) in FIG. 11 is formed in an arc shape separated by 120 degrees from each other in the cross section, and the space between the arc shapes is hollow. When this optical fiber structural material is actually used, the outer peripheral portion of the SUS frame 11 is covered with an annular outer layer 5 as shown in FIG. A pressure is applied as an external pressure to the SUS frame 11 through the outer layer 5.

In this case, the blood pressure that is the external pressure is the portion indicated by the cross section D1D1 of the SUS frame 11 in FIG. 9B as the pressure and the contact portion with the outer layer 5 and the SUS frame 11 (which can also be referred to as the structure 4). ), The SUS frame 11 is loaded on the fixed SM fiber, and axial strain occurs in the SM fiber. As a result, the frequency shift of the laser scattered light occurs in the SM fiber, and as a result, accurate blood pressure measurement becomes possible.
The structure 4 is a structural material that supports the rigidity of the entire catheter, and can be rotated by operating the operation handle 6 by connecting the operation handle 6 and the J tip.
Furthermore, it is also possible to implement a double-ended measurement method of scattered light by providing a reflection device such as FBG at the measurement end of the SM fiber tip. With such a configuration, it is possible to increase the measurement resolution of the Brillouin scattering frequency shift Δν _B , and it is possible to increase the accuracy of the calculated value based on the measurement data.

Embodiment 2. FIG.
As described above, in the first embodiment, desired data is obtained using the data of the frequency shift due to the reflected and scattered light of the two types of laser light measured with one SM fiber 3. Here, what obtains desired data by using only one type of scattered light will be described with reference to FIGS.

First, an example of a system configuration of the invention according to Embodiment 2 is shown in FIG. In the figure, an SM fiber 13 is further provided as a sensor fiber in addition to the SM fiber 3 shown in the first embodiment. In this case, the SM fiber 3 receives pressure load and is used as a pressure measurement fiber, while the SM fiber 13 is generally installed so as not to receive pressure load and is used as a temperature measurement fiber. . In the second embodiment, the SM fiber 3 is measured with the SM fiber 3 and the SM fiber 13 by utilizing the fact that the sensitivity coefficient related to the frequency shift of Rayleigh scattering is larger than the sensitivity coefficient related to the frequency shift due to Brillouin scattering. Among the frequency shifts generated in the laser light, only the signal of the Rayleigh scattering frequency shift Δν _R is used, and the frequency shift Δν _B due to Brillouin scattering is not used. At this time, the measurement by the SM fiber 13 may be performed simultaneously with the measurement of the frequency shift Δν _R of Rayleigh scattering by the SM fiber 3. Since other components are the same as those in the first embodiment, description thereof is omitted here.

Next, FIG. 11 is a detailed view showing a structure of a sensor portion of a catheter which is a blood pressure measurement probe equipped with the two SM fibers. In the figure, a SUS frame 11 having a hollow portion as a structure and having a rugby ball-like outer shape is provided around an SM fiber 3 (used for blood pressure measurement or the like) provided in a central axis portion. Further, an outer layer 5 that propagates the pressure of blood pressure as an external pressure and protects the SM fiber and the structure is provided on the outer peripheral portion so as to be in contact with the SUS frame at the maximum diameter portion of the SUS frame 11 that is the structure. (See FIGS. 11A and 11B). In addition to the SM fiber 3, it is free from pressure due to blood pressure as an external pressure, and is not fixed in the SM fiber axial direction and is free. Therefore, an SM fiber that can obtain a Rayleigh scattering frequency shift Δν _R reflecting only temperature changes. As shown in FIG. 11B, 13 is a position near the outer periphery of the catheter, and is arranged in the gap of the structure 4 so as not to contact the structure 4 and the outer layer 5.

Next, an analysis outline of the biodiagnostic sensor system according to Embodiment 2 will be described with reference to the flowchart of FIG. Since the optical fiber B detects only the change due to _the temperature change ΔT among _the Rayleigh scattering frequency shift Δν _R , when this frequency shift is expressed as Δν _R ^T , the sensitivity coefficient C ₂₂ of the optical fiber related to the temperature change is used. ΔT = Δν _R ^T / C ₂₂
On the other hand, when the frequency shift of Rayleigh scattering which is measured by the optical fiber A represents a Δν _R ^P, Δν _R ^P is, because it does not receive pressure, receives and elements that are affected by the distortion changes, the influence by the temperature changing element Therefore, Δν _R ^P = C ₂₁ Δε + C ₂₂ ΔT is established from Equation (4). Therefore, it expressed as ^{_{Δε = (Δν R P -C 22}} ΔT) / C 21.
In the measured object, the temperature change should be the same as the value measured with the optical fiber B and the value measured with the optical fiber A. By substituting the value of ΔT measured by the optical fiber B into an equation for obtaining Δε of the optical fiber A, Δε = (Δν _R ^P −Δν _R ^T ) / C ₂₁ (here, the optical fiber A and the optical fiber B). and sensitivity characteristics of the C ₂₂ will assume the same in). Since .DELTA..nu _R ^P and .DELTA..nu _R ^T are known quantities which are measured by an optical fiber, these from Motomari is [Delta] [epsilon], therefore, by using equation (2), P is determined as a function of [Delta] [epsilon] ( P = F (Δε)). Here, when the reference numerals shown in FIGS. 10 and 11 are used, the optical fiber A corresponds to the SM fiber 3, and the optical fiber B corresponds to the SM fiber 13.

As described above, even if only the frequency shift Δν _{R of} Rayleigh scattering is used, the SM fiber that is not affected by the pressure (blood pressure) is used as a means for measuring the frequency shift due to the temperature change, and a total of two SM fibers are used. By adopting the configuration, the same effects as in the first embodiment can be obtained also in the second embodiment. In the case of using a structure that continuously converts pressure into strain with high sensitivity in the above-described structure, the simplified structure of the present invention using a multipoint FBG is considered possible. However, the resolution is limited to the FBG interval, and the measurement distance is also limited.

  In the second embodiment described above, the configuration using two SM fibers as the sensor has been described, but the outer diameter of these two SM fibers is about 80 to 250 μm as described above, and is arranged on the outer periphery thereof. Even in consideration of the presence of the structure 4, a plurality of optical fiber sensors can be provided in the catheter. And even if another some optical fiber sensor is provided, the outer diameter of a catheter can be about 0.4 mm or less. An example of this is shown in FIG. In this figure, in addition to the two SM fibers 3 and 13, for example, two of an image fiber 14 for an endoscope and a fiber 15 for measuring a tomographic image (for example, optical coherence tomography (OCT)). An example in which an optical fiber is installed is shown. Instead of the image fiber 14 or the OCT measurement fiber 15, a projection agent charging tube may be installed. As described above, since a plurality of three or more optical fibers can be installed, multi-functional measurement can be performed.

Embodiment 3 FIG.
In actual PCI measurement and the like, there are cases where it is necessary to consider the influence on the blood pressure measurement due to the pulsation of the heart in addition to what has been described so far. In order to cope with such a case, the measurement system shown in the third embodiment is used. Specific examples will be described below with reference to the drawings.

  An example of the system configuration of the invention according to Embodiment 3 is shown in FIG. As shown in the figure, a pulsation sensing fiber 20 is connected to the measuring machine 8 in addition to the SM fiber 3 shown in the first embodiment. The heartbeat signal (see the center curve in FIG. 15) detected by the pulsation sensing fiber 20 is measured by the measuring device 8 in synchronization with the synchronization signal indicated by symbol M in FIG. To be recorded. In FIG. 15, the curve shown at the top is a model diagram showing a change example of the left ventricular pressure measured in synchronization with the synchronization signal, and it can be seen that the curve changes corresponding to the potential of the heartbeat signal. On the other hand, in the system shown in FIG. 14, data such as pressure (blood pressure) measured by the SM fiber 3 is also measured by the measuring device 8 in synchronization with the synchronization signal indicated by the symbol M in the same manner as the heartbeat signal. Recorded in the memory 9. Therefore, it is possible to compare both heartbeat signal data and pressure (blood pressure) data measured in synchronization with the synchronization signal indicated by symbol M on the same time axis, and by analyzing this, Data such as pressure (blood pressure) in consideration of the influence of pulsation on blood pressure measurement can be obtained. Specifically, for example, a plurality of blood pressure data with an improved S / N ratio can be obtained at a time by wrapping the pulsation sensing fiber 20 around the arm portion, sensing a plurality of pulsations caused by this, and synthesizing data based on this. And the effect of improving accuracy can be obtained.

  In FIG. 14, the SM fiber 13 shown in the second embodiment is not connected to the measuring machine, but the SM fiber 13 is added to the configuration shown in the second embodiment as the measurement system, that is, the SM fiber 3. In the measurement system in which the pulsation sensing fiber 20 described above is connected to the measuring device 8 in addition to the connected configuration, the blood pressure due to the pulsation of the heart is the same as described in the third embodiment. Data such as pressure (blood pressure) considering the influence on measurement can be obtained.

  Further, in the optical fiber type biodiagnostic sensor system according to the third embodiment, the data such as the time distribution of the pulsation peak at a predetermined position and the pulsation reproduction state based on the data obtained by the pulsation sensing fiber 20. Can be obtained as an analysis result with a synchronization signal level (for example, when the system shown in FIG. 14 is used, accuracy of about 4 msec (see FIG. 15)).

ここで、ρは血液の密度、ｒは血管の半径、ｔ_ｈは血管の肉厚である。この場合においては、実施の形態１の測定フローチャート中に示した表示例のように、２つの異なった時間帯でのデータを比較して、血管の動脈硬化の時間的変化のデータも得ることができるので、従来より有用なデータとなることが期待できる。Further, based on data such as time fluctuation of pulsation peak and pulsation reproduction state, data such as change in Young's modulus of blood vessels and data such as arrhythmia can be obtained as data used as a basis for biodiagnosis. For example, the elastic modulus data at a certain measurement time can be obtained by calculating the pulsation propagation velocity (also referred to as pulse wave velocity) v from the Maines-Corteweg equation (Non-patent Document 3) by the following equation (9). The Young's elastic modulus E indicating the hardness of the blood vessel can be derived, which is considered useful as data for diagnosing arteriosclerosis of blood vessels.

Here, [rho is the density of the blood, r is the radius of the vessel, is t _h a thickness of the blood vessel. In this case, as in the display example shown in the measurement flowchart of the first embodiment, the data at two different time zones can be compared to obtain data on temporal changes in vascular arteriosclerosis. Therefore, it can be expected to be more useful data than before.

  In the above description, the appropriate configuration of the above-described optical fiber structure material has been described with reference to FIG. 9 (hereinafter referred to as Example 1). It may be as shown in Example 2 or Example 3. It should be noted that the present invention can be freely combined with each other within the scope of the invention, and each embodiment can be appropriately modified or omitted.

(Example 2)
FIG. 16 shows another example of an appropriate configuration of the optical fiber structure material. The optical fiber deviates from the central axis of the measurement probe and is disposed in the immediate vicinity of the outer layer 5 near the outer periphery, and is fixed at the mandrel position fixed to the wire (FIG. 16A). And the outer layer 5 contacts an optical fiber with the load of the pressure which is a blood pressure. In this case, the load due to the pressure P is transmitted to the optical fiber through the outer layer 5, and as a result, the optical fiber is fixed and bent at the center position between the adjacent mandrels (see the broken line in FIG. 16B). In this case, the mandrel corresponds to the structure 4 that converts pressure into strain. FIG. 16C schematically shows the force relationship, and an axial force indicated by the symbol Tc is generated in the optical fiber.

(Example 3)
FIG. 17 shows still another example of an appropriate configuration of the optical fiber structure material. There are a total of three optical fibers indicated by solid lines, and one optical fiber is arranged in a 120-degree direction from each other as shown in the axial sectional view (see FIG. 17B). Further, these optical fibers are fixed to the position indicated by the symbol G1 in the axial direction by a fixing wire (see the left cross section G1G1 in FIG. 17B). Note that, at the position of the cross-section H1H1, the optical fiber is free without being constrained (see the diagram on the right side of FIG. 17B). The fixing wire is spirally wound on the cylindrical surface at a pitch Lp, and its outer shape forms a fixing wire outer envelope. Further, a driving wire is disposed on the central axis, and the optical fiber can be moved to a desired position. In this example, blood pressure as external pressure is transmitted to the optical fiber from the position where the fixing wire fixes the optical fiber. A structure (fixing wire) that is a protective layer that propagates the pressure of the object to be measured is provided on the outer peripheral portion.

1 catheter (blood vessel insertion type distributed pressure measuring device), 2 J tip,
3, 13 Single mode optical fiber (SM fiber), 4 structure,
5 outer layer, 6 operation handle, 7 multi-core SM fiber, 8 measuring machine,
9 data memory, 10 analysis / display device, 11 SUS frame,
12 adhesives, 14 image fibers,
15 fiber for tomographic image measurement (OCT fiber), 16 wire layer,
20 Pulsation sensing fiber.

Claims (10)

It is inserted into the blood vessel of a living body and measures the temperature and pressure distribution at a predetermined location of the measured object,
An outer layer that is deformed by external pressure and does not allow the object to be measured to enter the interior;
A single-mode optical fiber that is deformed by temperature and strain;
The single mode optical fiber is disposed so as to be in contact with the central axis portion or the outer peripheral portion, propagates the pressure applied to the outer layer, continuously converts the pressure distribution into the strain of the single mode optical fiber, and either fixed at a plurality of points in a part of the single-mode optical fiber, the elongation of the center axis of the part and proportionally said single mode optical fiber ambient pressure and contact at a plurality of locations of said single-mode optical fiber A structure to be converted, which supports the rigidity of the entire catheter having an outer diameter of 0.4 mm or less, and connects the catheter tip, which is a guide portion of the catheter, to the operation handle of the catheter, and rotates by the operation of the operation handle. Possible structures, and
A blood vessel insertion type distributed pressure measuring device comprising: The structure is a plurality of disk-shaped mandrels fixed to a wire shaft, and the single mode optical fiber is fixed to the outer periphery of the mandrel at a plurality of locations, and the single mode is applied by pressure load from the outer layer. 2. The blood vessel insertion type distributed pressure measuring device according to claim 1 , wherein an optical fiber bends in an inner circumferential direction between the adjacent mandrels. The structure is wound in a triple spiral shape, and the triple spirals are arranged at an outer peripheral position with a central angle of 120 degrees apart from each other in the axial cross section, and the wire outer envelope position in the axial direction. The blood vessel insertion type distributed pressure measuring device according to claim 1 , further comprising a fixing wire fixed to the outer periphery of the single mode optical fiber. The blood vessel insertion type distributed pressure measuring device according to claim 1 , wherein the structure has a hollow portion, is fixed to the single mode optical fiber at a central shaft portion, and is in contact with the outer layer at an outer peripheral portion. The blood vessel insertion type distributed pressure measuring device according to claim 1 , wherein multipoint measurement is performed using a multipoint FBG instead of the single mode optical fiber. The blood vessel insertion type distributed pressure measuring device according to claim 1, which measures the distribution of temperature and pressure at a predetermined location of the measurement object,
The laser beam is emitted to the single mode optical fiber, the frequency change of the scattered light generated in the single mode optical fiber is continuously detected, and the temperature of the single mode optical fiber obtained from the detected frequency change of the scattered light is detected. A measuring device that calculates and calculates a blood pressure at a predetermined position of the single-mode optical fiber from a change and a strain change;
A memory for storing a calculated value calculated by the measuring machine;
And an analysis / display device for performing a desired analysis or display based on the operation value stored in the memory,
An optical fiber type biomedical sensor system. 7. The optical fiber type biomedical sensor system according to claim 6 , wherein the measuring device detects a frequency change of Rayleigh scattered light and a frequency change of Brillouin scattered light. The blood vessel insertion type distributed pressure measuring device includes at least two single mode optical fibers, one of which is a single mode optical fiber that is configured to generate strain due to temperature change and not receive pressure, and The optical fiber type biomedical sensor system according to claim 6 , wherein the measuring device detects a frequency change of Rayleigh scattered light. 8. The optical fiber biodiagnostic sensor system according to claim 7 , wherein the single mode optical fiber has a reflection device at a terminal end. A pulsation sensing fiber that wraps around the arm and senses the heartbeat is detected. The pulsation sensing fiber detects the heartbeat signal, and controls the timing of detecting the frequency change of the scattered light in synchronization with the detected heartbeat signal. The optical fiber type biodiagnostic sensor system according to any one of claims 6 to 9 .

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